CN117015346A

CN117015346A - Dual layer detector system and method for spectral imaging and contrast enhanced digital breast tomosynthesis

Info

Publication number: CN117015346A
Application number: CN202280018696.3A
Authority: CN
Inventors: A·R·卢宾斯基; W·赵; 段晓雨; 黄海量; A·霍万斯基; J·斯塔夫罗; A·戈尔丹
Original assignee: Research Foundation of State University of New York
Current assignee: Research Foundation of State University of New York
Priority date: 2021-03-01
Filing date: 2022-03-01
Publication date: 2023-11-07
Also published as: AU2022230968A1; WO2022187188A1; KR20230153376A; US20240156420A1; JP2024514393A; EP4301232A1; CA3210335A1

Abstract

Structures and methods operable to detect radiation are described. The structure includes a dual layer detector imaging device that allows dual energy imaging of single shot x-rays. In one embodiment, the front layer of the detector comprises a photon counting detector and the back layer of the detector comprises an energy integrating detector comprising a scintillating phosphor screen. In further embodiments, the front detector layer comprises a direct conversion x-ray detector and the second rear detector layer comprises an indirect conversion x-ray detector. The imaging system also includes a spatial separation filter positioned adjacent the x-ray radiation source for absorbing x-ray radiation to separate the radiation into low energy and high energy components for incidence on the imaging subject. In an embodiment, the imaging subject comprises a contrast material having a characteristic K-edge atomic band level, and the separation filter absorbs X-ray radiation in the vicinity of the K-edge atomic band level.

Description

Dual layer detector system and method for spectral imaging and contrast enhanced digital breast tomosynthesis

Cross Reference to Related Applications

The application claims the benefit of U.S. provisional application No. 63/154,879, filed on 3/1/2021, the entire contents of which are incorporated herein by reference.

Technical Field

The present application relates generally to radiation detectors and digital radiography and 3D digital breast imaging.

Background

In digital radiography, an imaging system may include a Flat Panel Detector (FPD) comprising a collection of layers, such as a scintillator screen that absorbs radiation and produces visible light pulses upon absorption of x-rays, a pixelated array of photosensors (e.g., photodiodes) in which the produced light is sensed, and a thin film transistor array for generating electrical signals. The generated electrical signals may be used by the imaging system to produce a digital image. In some examples, the quality (e.g., sharpness, resolution) of the generated image may be affected by various phenomena, such as light scattering and/or other phenomena.

Spectral x-ray imaging extracts an image of a specific material of the volume of interest. This technique is used in clinical radiography to provide additional diagnostic information about the anatomy of the patient. Spectral imaging is performed by: acquiring two or more x-ray transmission images (i.e. "projections") of a volume of interest, each image having a different x-ray energy; and applying post-processing techniques to identify the constituent material of the volume by differences in x-ray attenuation characteristics of the constituent material of the volume. One strategy to achieve energy separation between images is through time subtraction, where the x-ray source kVp (kilovoltage peak) and/or filter is changed between successive x-ray projections

Alvarez, in US4, 029,963, describes improving contrast or signal-to-noise ratio (SNR) of material selective x-ray images by using high-energy and low-energy projection measurements. US4,445,226 to Brody discloses a method for eliminating soft tissue or bone structures from x-ray images using a hybrid energy subtraction technique. US8,792, 617 to Baetz discloses a method of creating dual energy x-ray images in mammography using two exposures, each having a different kVp and filtering. However, this approach has problems and artifacts due to the patient's movement between exposures.

Fig. 1B depicts a prior art solution 10 for a dual energy subtraction, post-contrast injection or time subtraction strategy before and after contrast injection, which has been used in dual shot/single detector spectral mammography systems and in spectral Computed Tomography (CT) systems. This approach uses fast kV switching to produce alternating high-energy and low-energy projections. For example, in the dual shot approach of fig. 1B, the first filter 20 is located near the front of the X-ray radiation source 15 for passing only low energy X-ray radiation 25 incident on the imaging subject 12. A single detector 30, such as an X-ray photoconductor or a scintillation screen, receives photons from the low energy X-ray radiation 25 that are transmitted through the object 12 and are not absorbed by the object 12, and associated circuitry converts these detected photons into electrical signals for producing a first Low Energy (LE) image of the object. Subsequently, in the dual shot approach of fig. 1B, a second filter 40 is located in front of and near the X-ray radiation source 15 for passing only the high energy X-ray radiation 25 incident on the imaging subject 12. The single detector 30 receives photons from the high energy X-ray radiation 45 that are transmitted through the object 12 and are not absorbed by the object 12, and associated circuitry converts these detected photons into electrical signals for producing a first High Energy (HE) image of the object. The two shot approach of performing the two exposures can achieve a high degree of spectral separation between the high and low energy images because they allow for varying the x-ray source kVp, filters, and image receiver in the acquisition workflow. While this flexibility is desirable for achieving high contrast spectral imaging, this approach is inherently limited by misregistration artifacts caused by subject motion between acquisitions.

Alternatively, US9,526,466 to Karim discloses a method of producing dual energy x-ray images in mammography using stacked integrated multi-layer detectors, wherein the detectors are simultaneously exposed to an x-ray beam in the same x-ray exposure. However, due to the lack of a spectral separation filter as in the present disclosure, this approach suffers from limited energy separation between the low-energy and high-energy images produced, and results in loss of contrast and SNR in the energy subtraction image.

Finally, danielsson's U.S. patent 7,342,233 claims a device with an array of photon counting channels, wherein each channel converts a separate x-ray detection event into electrical pulses according to pulse height, and wherein each of the two sets of counters counts pulses according to whether the pulses are above or below a given threshold to produce high and low energy images in a single exposure.

Disclosure of Invention

Accordingly, structures, imaging systems, and detectors are disclosed that provide improved image quality and dose performance. The system and imaging method include a single or dual layer detector that allows one shot (single exposure) x-rays to be used for dual energy imaging, i.e., single shot energy discrimination.

In an example embodiment, a system comprising a single or dual layer detector includes a spectral separation filter positioned near the output of an X-ray energy source for modulating X-ray radiation into low and high energy bands to allow one shot of X-rays for dual energy imaging of an object.

In an embodiment, the imaging subject may include a contrast material having a characteristic K-edge atomic energy band level. The separation filter is made of a material that absorbs X-ray radiation near the K-edge atomic band level to produce two X-ray radiation bands: low Energy (LE) radiation bands and High Energy (HE) radiation bands for dual energy imaging of an object.

The apparatus may further include a single layer X-ray energy image detector including a pre-X-ray imaging Photon Counting Detector (PCD) for receiving incident X-ray radiation transmitted through the spectral separation filter and the imaging subject and generating an electrical signal capable of producing a first LE image of the imaging subject and generating an additional electrical signal capable of producing a second HE image of the imaging subject.

In an embodiment, the PCD is an amorphous selenium (α -Se) x-ray photon counting flat panel imager (SWAD) for detecting both lower energy and higher energy photons to form respective low energy and high energy images.

In addition, the apparatus may include a spectral separation filter and a dual layer X-ray energy image detector including a first front direct conversion X-ray imaging detector located on an underlying substrate for generating LE and HE images of the subject and a second rear indirect conversion X-ray imaging detector located below the substrate for generating information to be combined with the HE images of the subject.

In this embodiment, the first front X-ray imaging detector may comprise a PCD, such as an amorphous selenium (α -Se) X-ray photon counting flat panel imager (SWAD) for detecting both lower energy and higher energy photons used to form respective low energy and high energy images. The second post X-ray imaging detector comprises an indirect conversion flat panel X-ray detector or a photon energy integration detector and comprises a material having an atomic number selected to effectively detect higher energy photons to form a high energy image for enhancing the higher energy image obtained by the first front PCD.

In yet a further embodiment of the dual layer X-ray imaging detector, both the first front X-ray imaging detector and the second rear X-ray imaging detector comprise integral detectors. The front-integrating detector includes a first pixel sensor for directly converting first energy band photons of incident radiation transmitted through the imaging subject into a first image signal configurable to form a low-energy image of the imaging subject. The post-integration detector formed below the substrate includes a second pixel sensor for converting second energy band photons of incident radiation transmitted through the imaging subject and the pre-integration detector and the substrate into a second image signal configurable to form a high energy image of the imaging subject.

According to a first aspect of the present invention there is provided an apparatus comprising: a separation filter for separating a radiation spectrum from the X-ray radiation source into a first energy band and a second energy band for imaging incident radiation on the object; a substrate; an x-ray photon counting detector formed on the substrate, the x-ray photon counting detector comprising an array of detector pixels, each detector pixel comprising a sensor for detecting interaction of individual x-ray photons of incident radiation transmitted through the imaging subject during a fixed period of time; and each detector pixel of the array has an associated counting circuit operable to generate a respective counted first electrical signal representative of the number of detected interactions of the respective x-ray photons of the first energy level band and a respective counted second electrical signal representative of the number of detected interactions of the respective x-ray photons of the second energy level band, wherein the first electrical signal and the second electrical signal from the detector pixels of the array provide respective energy spectral images of the imaging subject.

In a further aspect, an apparatus comprises: a separation filter for separating a radiation spectrum from the X-ray radiation source into a first energy band and a second energy band for imaging incident radiation on the object; a first substrate; an x-ray photon pre-count detector formed on the first substrate, the x-ray photon count detector comprising an array of detector pixels, each detector pixel comprising a sensor for detecting interaction of individual x-ray photons of incident radiation transmitted through the imaging subject during a fixed period of time; each detector pixel of the array has an associated counting circuit operable to generate a respective counted first electrical signal representative of a number of detected interactions of individual x-ray photons of a first energy band and a respective counted second electrical signal representative of a number of detected interactions of individual x-ray photons of a second energy band, wherein the first and second electrical signals from the detector pixels of the array provide respective energy spectral images of the imaging subject; and a rear detector formed on the second substrate and located under the first substrate, the rear detector comprising: a scintillation screen for converting incident radiation comprising x-ray photons transmitted through the imaging subject and through a second energy band of the front detector into light photons; and a photosensor array disposed between the scintillation screen and the second substrate for the rear detector, the photosensor array being operable to capture photons of light from the scintillation screen and to convert the captured photons of light into additional electrical signals that are operable to be combined with the second electrical signals from the front detector pixel array to obtain an image of the imaging subject.

In a further aspect, an apparatus comprises: a separation filter for spectrally separating radiation from the X-ray radiation source into a first energy band and a second energy band for imaging incident radiation on the object, the second energy band having an energy greater than the first energy band; a glass substrate; a front-integrating detector formed on the glass substrate, the front-integrating detector comprising a first pixel sensor for directly converting first energy band x-ray photons of incident radiation transmitted through the imaging into a first image signal configurable to form a low-energy image of the imaging subject; and a post-integration detector formed below the glass substrate, the post-integration detector including a second pixel sensor for indirectly converting second energy band x-ray photons of incident radiation transmitted through the imaging subject and through the pre-integration detector and the substrate into a second image signal configurable to form a high energy image of the imaging subject.

Further to this aspect, the front integration detector includes: a first photoconductive layer for converting incident radiation comprising x-ray photons transmitted through a first energy band of the imaging subject into electrical charge; and a first charge storage array disposed between the first photoconductive layer and the substrate for storing charge associated with the converted x-ray photons.

Further to this aspect, the post-integration detector includes: a scintillation screen for converting incident radiation comprising x-ray photons transmitted through a second energy band of the imaging subject into light photons; and a photosensor array disposed between the second scintillation screen and the substrate, the photosensor array being operable to capture photons of light from the second scintillation screen and to convert the captured photons of light into a second imaging signal.

In this aspect, the imaging subject includes a contrast material having a characteristic K-edge atomic energy band level, and the separation filter has an X-ray absorption edge for absorbing X-ray radiation near (e.g., within 10 keV) the K-edge atomic energy band level of the contrast material.

Further features, as well as the structure and operation of various embodiments, are described in detail below with reference to the accompanying drawings. In the drawings, like reference numbers indicate identical or functionally similar elements.

Drawings

FIG. 1A illustrates an example dual layer detector imaging system in one aspect of the present disclosure that allows for single shot x-rays for dual energy imaging.

FIG. 1B depicts a standard single detector system that requires two shot x-ray exposures for dual energy imaging;

Fig. 2A-2E show example graphs depicting a comparison of HE and LE spectral separations for a dual shot/single detector and dual layer detector scheme.

FIG. 2F depicts a table summarizing example average LE and average HE spectral energy bands obtained from each of the embodiments of FIGS. 2A-2E;

3A-3C illustrate the effect of filter thickness and thickness of top detector material (T1) on spectral separation achieved by the dual detector spectral imaging scheme depicted in the embodiment of FIG. 1A; FIG. 3D is a table summarizing the improvement in image quality measured by SDNR for the single shot/dual layer detector of FIG. 1A for front and rear detector layers of different thickness;

FIG. 4 shows a first embodiment of a spectral imaging system comprising a single shot/single layer detector for producing dual LE and HE spectral images of an object;

FIG. 5 depicts a cross-sectional schematic diagram of a fully assembled SWAD device for use as a front detector in a single detector or dual detector spectral imaging scheme according to an embodiment;

FIG. 6 depicts example results of a simulated pixel response of a photon counting SWAD device such as that depicted in FIG. 5;

FIG. 7 depicts a second embodiment of a spectral imaging system that includes a single shot/dual layer detector for producing dual LE and HE spectral images of an object;

FIG. 8 depicts a further embodiment of a spectral imaging system including a single shot/dual layer detector for producing dual LE and HE spectral images of a subject;

FIG. 9 illustrates a further embodiment of a spectral imaging system including a single shot/dual layer detector including a direct conversion x-ray front x-ray detector layer and an indirect conversion rear x-ray detector layer x-ray detector;

FIG. 10 illustrates a further embodiment of a spectral imaging system in aspects of the present disclosure, including a dual shot exposure/dual layer detector with a source array or a step imaging tomography system;

FIGS. 11A-11E depict simulated result images of digital mammography, contrast-enhanced dual-energy mammography, digital breast tomosynthesis, and contrast-enhanced digital breast tomosynthesis;

FIG. 12 shows a table depicting signal difference to noise ratios in the simulated contrast enhanced dual energy mammography results of FIGS. 11C, 11D, and 11E; and

fig. 13A-13C depict respective LE DBT image slices, HE DBT image slices, and CE DBT image slices, which illustrate the problems to be overcome by prior art dual shot imaging schemes.

Detailed Description

Aspects of the present disclosure will be described in detail below with reference to the accompanying drawings. In the present disclosure, for clarity of understanding of the concepts of the present disclosure, explanations concerning related functions or constructions known in the prior art have been omitted to avoid obscuring the present disclosure in unnecessary detail.

Spectral X-ray imaging extracts a specific material image of the volume of interest. This technique is used in clinical radiography to provide additional diagnostic information about the anatomy of the patient. Spectral imaging is performed by: acquiring two or more x-ray transmission images (i.e. "projections") of a volume of interest, each image having a different x-ray energy; and applying post-processing techniques to identify the constituent material of the volume by differences in x-ray attenuation characteristics of the constituent material of the volume.

In the embodiments herein, there are two strategies to achieve energy separation between images: 1. time subtraction, where the x-ray source kVp and/or filter changes between consecutive x-ray projections, or 2. Single shot energy discrimination in which multiple detectors sensitive to different energy spectra are provided. In embodiments herein, single shot energy discrimination is performed using a single Photon Counting Detector (PCD), where two or more energy bins are used to form multiple specific energy images from a single x-ray projection. In further embodiments, single shot energy discrimination is performed using a dual detector comprising a photon counting detector and a photon Energy Integration Detector (EID) for forming a plurality of specific energy images from a single x-ray projection. Further embodiments enable single shot energy discrimination using a dual detector comprising a two photon EID to form multiple specific energy images from a single x-ray projection. Further embodiments contemplate dual detectors using two or more radiation exposures (multiple shots) to form multiple specific energy images. These two strategies may be combined to further enhance the energy discrimination of spectral imaging.

Spectral imaging using single shot energy discrimination is free of motion misregistration artifacts compared to conventional 2D Full Field Digital Mammography (FFDM), spectral FFDM, and conventional DBT, and as described further herein, can be used in spectral DBT applications to provide more complete and accurate diagnostic information about the breast. This information includes 3D breast tissue density, 3D microcalcification distribution and type, 3D distribution of contrast agent (e.g., iodine), and 3D material decomposition of breast lesions of interest (e.g., masses). The present disclosure describes a spectral imaging system implementation that may be used to provide one or all of the above information, for example, in a spectral DBT.

Fig. 1A illustrates one embodiment of a dual layer detector spectral imaging system 100 that allows single shot x-rays to be used for dual energy imaging of a subject 12.

In an embodiment, the spectral imaging system 100 includes an X-ray source 15, which may be an X-ray tube that generates X-rays, or other device that may generate X-rays. The X-ray source may irradiate X-ray radiation onto the body through a spectral separation filter, wherein the body may absorb a portion of the X-rays, resulting in X-ray attenuation. Attenuated X-rays may be directed toward dual detector structure 101 as incident X-rays 122.

In either embodiment, the dual layer X-ray spectral imaging detector arrangement includes an X-ray filter 120 positioned near the front of the X-ray radiation source 15 for absorbing a portion of the X-ray radiation output of the radiation source 15 such that the X-ray radiation 122 passes simultaneously at two energy levels, a first LE energy level including a radiation energy band below the X-ray absorption edge of the separation filter and a second HE energy level including a radiation energy band above the X-ray absorption edge of the separation filter. The system 100 includes a dual layer X-ray imaging detector for receiving incident X-ray radiation 122 transmitted through the subject 12. The dual layer X-ray imaging detector is configured as a stack 101 comprising a first front detector 110 disposed on a substrate 125 and a second rear detector 150 located or attached under the substrate 125 using, for example, a radiation transparent adhesive. In an embodiment, front detector 110 may include a Photon Counting Detector (PCD) or a photoconductive-type photon integrating detector. The post detector 150 may comprise a light photon integrating detector. In these embodiments, the substrate may include a substrate 125 of glass or similar material.

In some embodiments herein, the spectral separation filter 120 at the output of the x-ray source includes an atomic number Z _F1 To Z _FN Thickness T _F1 To T _FN To modulate the x-ray energy spectrum incident on imaging subject 12. A filter material having an atomic number is selected that selectively removes x-ray energy from the beam by preferential attenuation of energy near its K-edge, e.g., energy within 10 keV. The filter materials, thicknesses, and their order of arrangement relative to the x-ray source are selected to shape the energy transmission characteristics of the filter based on the initial energy spectrum, detector characteristics, and spectral information of interest. For example, a material may be selected that has a K-edge near (e.g., within 10 keV) to the contrast material (e.g., iodine) of the breast imaging. Example separation filter materials may include rhodium, silver, palladium, indium, and tin filters.

In an embodiment, dual layer detector imaging system 100 is particularly configured for Digital Breast Tomosynthesis (DBT), and is particularly useful for obtaining high resolution three-dimensional (3D) X-ray images of breast 12. In such embodiments, the first front detector 110 is a direct conversion flat panel x-ray detector ("front detector") that is first exposed to the x-ray beam transmitted through the imaging subject, having a lower atomic number (e.g., atomic number Z ₁ ) And thickness T ₁ And high spatial resolution. The atomic number and thickness of the pre-detector material are selected to preferentially absorb lower energy x-rays while transmitting higher energy, thereby allowing a low energy image to be formed. Its high spatial resolution is used to preserve image detail information, i.e. small structures and sharp edges. In an example, the front detector 110 is an amorphous selenium detector (e.g., Z ₁ ＝34，T ₁ =150 microns or 200 microns).

In an embodiment, the second post-detector 150 is of atomic number Z ₃ ＞Z ₁ And thickness T ₃ Is an indirect conversion flat panel x-ray detector ("post detector"). The atomic number of the post-detector is selected to effectively detect higher energy photons, thereby forming a high energy image. The atomic number of the post-detector can be matched with the K-edge of the contrast agentCsI of a ligand, e.g., iodine: tl to increase the saliency of the contrast agent in the high-energy image. The post-detector may include a scintillator that may be transparent or optically turbid, structured or unstructured (e.g., columnar Cs: tl or powder Gd ₂ O ₂ S: tb), and includes an optically reflective or absorptive backing. The scintillator is coupled to the photodetector array by direct deposition, pressure contact, or in some embodiments by using a fiber optic plate to transmit the x-ray induced light image to the photodetector without light diffusion. The light sensor may be a-Si: h-photodiodes, MIS types, or other types known in the art. The Thin Film Transistor (TFT) switching element may be a-Si: type H, metal Oxide (MOTFT) or other types known in the art

Fig. 2A-2E illustrate various graphs depicting the ability of the dual layer scheme 100 of fig. 1A to spectrally separate the x-ray spectrum into LE and HE spectral energy bands for incidence on an imaging subject. For example, each of the graphs of fig. 2A-2D shows a spectral separation of x-ray radiation, and in particular, the exemplary graph depicts the number of photons received in the filtered radiation forming LE and HE spectral imaging bands as a function of photon energy (in keV) of the filtered radiation, as implemented by the dual layer detector scheme 100 of fig. 1A for the W/Rh spectrum. A comparison plot of the spectral separation of x-ray radiation is additionally shown, and in particular, an exemplary graph depicts the number of photons received in the filtered radiation as a function of photon energy (in keV) of the filtered radiation, as achieved by the dual shot/single detector scheme 10 of fig. 1B. In the example graphs shown in fig. 2A-2D, the detector is a 150 micron thick amorphous selenium (α -Se) Energy Integrating Detector (EID), and the spectral separation filter includes 200 micron thick rhodium.

In a first depiction, fig. 2A shows a comparison 201 of spectral separations achieved by the single shot/dual layer detector scheme of fig. 1A as compared to the dual shot/single layer detector approach of fig. 1B. As shown in fig. 2A, LE spectral radiant energy 214 detected by the single shot/dual layer detector scheme far exceeds LE spectral radiant energy 204 detected by the dual shot/single layer detector scheme of fig. 1B. Similarly, as shown in the graph of fig. 2A, the single shot/dual layer detector scheme detects HE spectral radiant energy 216 that exceeds the HE spectral radiant energy 206 detected by the dual shot/single layer detector scheme of fig. 1B. Furthermore, LE and HE radiation achieves a widened separation 215 of about 20keV, resulting in improved LE and HE images.

Fig. 2B shows a comparison 221 of spectral separation achieved by the single shot/dual layer detector scheme of fig. 1A compared to the single layer detector scheme of fig. 1B, the single shot/dual layer detector scheme of fig. 1A using a k-edge band spectral separation filter for single shot LE/HE radiation separation, and the single layer detector scheme of fig. 1B using a non-k-edge filter for LE and HE spectral separation. As shown in fig. 2B, the single shot/dual layer detector scheme using a k-edge band filter detects LE spectral radiant energy 234 and HE spectral energy 236, each having a spectral energy peak separation 225 of about 20keV, which far exceeds the spectral separation achieved by the dual shot/single layer detector and non-k-edge filter scheme of fig. 1B, which is shown in fig. 2B as LE spectral radiant energy 224 and HE spectral radiant energy 226 with little or no spectral energy separation.

Fig. 2C shows a graph 241 of spectral separation achieved for the single shot/dual layer detector scheme of fig. 1A using the front alpha-Se photoconductive detector 110, the rear CsI energy-integrating scintillation detector 150, and the k-edge band spectral separation filter for single shot LE/HE radiation separation. As shown in fig. 2C, a single shot/dual layer detector scheme using a front alpha-Se photon integrating detector 110, a rear photon integrating detector 150, and a k-edge band filter detects LE spectral radiant energy 244 and HE spectral energy 246. Enhanced spectral energy peak separation 245 of greater than 20keV is achieved between LE and HE bands, resulting in improved LE and HE images.

Fig. 2D shows a graph 261 of spectral separation achieved for the single shot/dual layer detector scheme of fig. 1A using the front alpha-Se photoconductive detector 110, the rear alpha-Se photoconductive detector 150, and the k-edge band spectral separation filter for single shot LE/HE radiation separation. As shown in fig. 2D, a single shot/dual layer detector scheme using a front alpha Se photon integrating detector 110, a rear alpha Se photon integrating detector 150, and a k-edge band filter detects LE spectral radiant energy 254 and HE spectral energy 256. Enhanced spectral energy peak separation 255 of greater than 20keV is achieved between LE and HE bands, resulting in improved LE and HE images.

Fig. 2E shows a graph 281 of spectral separation implemented for the single shot/dual layer detector scheme of fig. 1A, which uses the front CsI energy integrating scintillation detector 110, the back CsI energy integrating scintillation detector 150, and the k-edge band spectral separation filter for single shot LE/HE radiation separation. As shown in fig. 2E, a single shot/dual layer detector scheme using front CsI photon integrating detector 110, back CsI photon integrating detector 150, and k-edge band filters detects a large amount of LE spectral radiant energy 274 and HE spectral energy 276.

Fig. 2F is a table 290 depicting a column 292 providing an average of the detected LE spectral band energies in keV and a column 294 providing an average of the detected HE spectral band energies in keV for the embodiment of fig. 2A-2E. As shown in fig. 2F, while an average LE and average HE spectral separation of about 17keV is achieved for the dual shot/single layer detector, the dual layer detector configuration of fig. 1A with k-edge band filtering and the front a-Se photoconductive detector 110 achieves a comparable average LE and average HE band separation of about 13 keV.

Figures 3A-3C illustrate the effect of filter thickness and top detector material thickness (T1) on spectral separation achieved by the dual detector spectral imaging scheme depicted in the embodiment of figure 1A.

Fig. 3A particularly shows a comparison 300 of LE and HE spectral energy separations implemented for the single shot/dual layer detector scheme of fig. 1A, which uses two different thicknesses of rhodium spectral separation filter 120. In particular, as shown in fig. 3A, a single shot/dual layer detector scheme using a 200 micron thick Rh k-sideband filter detects LE spectral radiant energy 304 and HE spectral energy 306 with a corresponding spectral energy peak separation 315 of about 20keV, resulting in improved LE and HE images. For comparison, fig. 3A depicts LE spectral radiant energy 305 and HE spectral radiant energy 307 implemented for a single shot/dual layer detector scheme using a 100 μm Rh k edge band filter. As shown in fig. 3A, approximately the same spectral energy separation is achieved in this single shot/dual layer detector scheme.

Fig. 3B particularly shows a comparison 320 of LE and HE spectral energy separations implemented for the single shot/dual layer detector scheme of fig. 1A, which uses two different thicknesses of the front a-Se photoconductive detector 110 and a 200 micron thick Rh spatial separation filter. In particular, as shown in fig. 3B, LE spectral radiant energy 314 and HE spectral energy 316, each having a spectral energy peak separation 325 of about 20keV, are detected using a single shot/dual layer detector scheme of the front alpha-Se photon integrating detector 110 of 200 microns thickness, resulting in improved LE and HE images. For comparison, fig. 3B depicts LE spectral radiant energy 317 and HE spectral radiant energy 319 implemented for a single shot/dual layer detector scheme of the 150 micron thick front alpha-Se detector 110. As shown in fig. 3B, approximately the same spectral energy separation is achieved in this single shot/dual layer detector scheme.

Fig. 3C illustrates the effect of providing a filter material between the front detector layer 110 and the rear detector layer 150 on the single shot/dual layer detector scheme of fig. 1A. Fig. 3C particularly shows a comparison 350 of various graphs of the quantum detection efficiency measurements of the detector as a function of detected photon energy levels detected for different glass substrate thicknesses, e.g., glass substrate thickness of the front detector ranging between 300 microns-700 microns, optionally with the addition of a glass substrate 352 from the bottom detector (when it is in the back-illuminated geometry). In contrast, an optional copper filter (or other higher z material) may be used.

Fig. 3D is a table 375 summarizing the resulting improvement in image quality measured by the signal-to-noise ratio (sdn r) 380 for different thicknesses of the front detector layer 110 and the rear detector layer 150 of the single shot/dual layer detector of fig. 1A. In the first row, a single shot/dual layer detector with a front alpha-Se detector 110 of about 200 microns thickness and a rear CsI photon integrating detector 150 of about 400 microns thickness and a Rh filter of about 200 microns thickness allows detection of an average LE spectral radiant energy of 27.5keV and an average HE spectral energy of about 40.9keV, providing an improved SDNR of about 6.01 at an X-ray tube source load of about 108.2 mAs. In the second row, a single shot/dual layer detector with a front alpha-Se detector 110 of about 200 microns thickness and a rear CsI photon integrating detector 150 of about 400 microns thickness and a Rh filter of about 100 microns thickness allows detection of an average LE spectral radiant energy of 28.0keV and an average HE spectral energy of about 39.6keV, providing an improved SDNR of about 5.22 at an X-ray tube source load of about 18.2 mAs. Similarly, a single shot/dual layer detector having a front alpha-Se detector 110 of about 150 microns thickness and a rear CsI photon integrating detector 150 of about 400 microns thickness and a Rh filter of about 200 microns thickness allows for detection of an average LE spectral radiant energy of 26.8keV and an average HE spectral energy of about 39.9keV, providing an improved SDNR of about 6.31 at an X-ray tube source load of about 108.2 mAs.

Several embodiments of the dual layer detector spectral imaging system scheme of fig. 1A for imaging a subject 12 are provided, each specifically configured for Digital Breast Tomosynthesis (DBT), i.e., 3D breast spectral imaging. The 3D spectral x-ray imaging system includes various combinations of six components: 1) Spectral separation filter at the output of an x-ray source comprising one or more atomic numbers Z _F1 To Z _FN And thickness T _F1 To T _FN To modulate the x-ray energy spectrum incident on the imaging subject; 2) Direct conversion flat panel x-ray detector ("front detector") which is first exposed to an x-ray beam transmitted through the imaging subject, has a lower atomic number Z ₁ Thickness T ₁ And high spatial resolution; 3) A filter at the exit surface of the front detector comprising an atomic number Z _2F To Z _NF And has a thickness T _2F To T _2F Modulating the x-ray energy spectrum exiting the front detector; 4) A second indirectly-converted flat-panel x-ray detector ("post detector") at the filter exit surface, withHaving atomic number Z ₃ ＞Z ₁ And thickness T ₃ The method comprises the steps of carrying out a first treatment on the surface of the 5) A combined dual layer detector with dual shot x-ray exposure. This can be achieved by a switchable kVp and filtered DBT system that provides up to four images for multiple material decomposition; and 6) Photon Counting Detector (PCD) versions of alpha-Se flat panel imagers (e.g., field shaping multi-well avalanche detectors or SWADs) as top (front) detectors.

Fig. 4 shows a first embodiment of a spectral imaging system comprising a single shot/single layer detector 400. This first embodiment comprises a combination of spectral separation filters 120, the spectral separation filters 120 being located in the vicinity of the front of the X-ray radiation source 15 for absorbing a part of the X-ray radiation output of the radiation source 15. The spectral imaging system includes a single direct conversion flat panel x-ray detector, such as a Photon Counting Detection (PCD) device 401, positioned on a substrate, such as a glass substrate 225. As described in more detail below, PCD 401 includes a flat panel imager, such as an amorphous selenium (a-Se) based field shaping multi-well avalanche detector (SWAD), for photon counting at each pixel location. The alpha-Se x-ray photon counting flat panel SWAD imager layer 401 converts the received photons into electrical pulses having pulse heights commensurate with the energy levels of the received photons. At each pixel location of the PCD layer, the semiconductor threshold and count circuit 230 detects the pulse height to distinguish any one of the LE photons and HE photons received at the PCD layer from incident radiation that is not absorbed by (i.e., transmitted through) the object 12. The threshold circuit may include a first threshold circuit for rejecting the noise signal and a second pulse threshold circuit for comparing the height of the received pulse to a pulse height threshold. At a pixel, LE and HE counter circuits count the number of photon interactions as an electronic pulse received at that pixel location. Photons received from radiation that exceed a pulse height threshold may be counted as HE photons, and the HE stores a "bin" maintaining a photon count value in a memory store associated with the pixel location. Similarly, at the same pixel location, photons received from radiation below the pulse height threshold may be counted as LE photons, and the LE "bin" maintains an LE photon count value in a memory store associated with that pixel location.

As shown in fig. 4, a data acquisition circuit 98 including a programmed processor 99 and memory configured to communicate with each other may receive the count values from each of the LE and HE bins and process the values to create respective LE and HE spectral images of the subject.

For embodiments employing a Photon Counting Detector (PCD) version of an alpha-Se x-ray photon counting flat panel imager (SWAD) as the top (front) detector, fig. 5 depicts a schematic cross-sectional view of a fully assembled SWAD device. In the cross-sectional schematic of the fully assembled SWAD device 500 of fig. 5, two distinct regions are depicted: (1) A body region 510 formed of an alpha-Se thick film for x-ray absorption, and (2) a multi-well sensing region 525 composed of a plurality of wells 526a, a plurality of pillars 550 deposited directly on the readout electronics, such as the semiconductor pixel threshold and CMOS count circuit 230 of fig. 4. The sidewalls of the well 526 are formed by a dielectric pillar 550 having two gate electrodes 530, 535 embedded therein, the gate electrodes 530, 535 being operable to generate a local high field within the well region for tunable avalanche gain and electrostatic shielding of the pixel electrodes for fast unipolar time differential charge sensing. When the gate electrodes 530, 535 are biased, the field lines over the wells bend, creating a field shaping effect that directs drift carriers into the wells. The electric field strength within the well can be tuned to create a localized high field region in which carriers can experience avalanche gain while maintaining a low field directly above the readout electronics and throughout the body. Furthermore, the embedded gate electrode acts as a Frisch gate, electrostatically shielding the pixel electrode from sensing any carrier movement within the body, while creating a strong near field effect, where the collected signal increases rapidly as carriers enter the well and experience avalanche gain. The body region (200 microns) is an order of magnitude thicker than the well region to provide good x-ray absorption and minimize the effect of depth dependent avalanche gain. The common high voltage electrode 560 is positively biased so that fast carriers (holes, in a-Se) can be collected.

Fig. 6 depicts the results 600 of a simulated pixel response of a photon counting SWAD device such as depicted in fig. 5, wherein a 300 micron thick alpha-Se body layer was exposed to a 49kVp W spectrum filtered with 237 micron copper, and a 4cm thick breast had a 50%/50% fibroglandular/adipose tissue composition. Pixel responses for both 75 micron and 100 micron pixel sizes are shown. The simulation result 600 is based on a multi-pixel geometry whereby the pixel response shown includes spectra collected by a single center pixel and crosstalk from 8 adjacent pixels, combined with the charge-sharing spatial energy effect.

In an embodiment, using a photon counting version of an α -Se x-ray photon counting flat panel imager (SWAD) as the top detector provides a low cost alternative to other Photon Counting Detectors (PCD) using crystalline Cd (Zn) Te. Its energy resolution and count rate depend on the geometry of the Frische gate built on top of the CMOS photon counting integrated circuit. With an avalanche gain of 10 and a linear Frische gate, a count rate of 100k counts per second (cps) is possible with an energy resolution of 3keV.

Fig. 7 depicts a second embodiment of a spectral imaging system that includes a single shot/dual layer detector 700. This embodiment includes a spectral separation filter 120 located near the front of the X-ray tube source 15 for absorbing a portion of the X-ray radiation output of the source 15.

In an embodiment, the spectral separation filter 120 at the output of the x-ray source comprises one or more materials, e.g., having an atomic number Z _F1 To Z _FN And thickness T _F1 To T _FN To modulate an x-ray energy spectrum 122 incident on the imaging subject 12. A filter material having an atomic number is selected that selectively removes x-ray energy from the beam by preferential attenuation of energy near its K-edge, e.g., energy within 10 keV. The filter materials, thicknesses, and their order of arrangement relative to the x-ray source are selected to shape the energy transmission characteristics of the filter based on the initial energy spectrum, detector characteristics, and spectral information of interest. For example, a material may be selected where the K-edge is close to the breast imaging contrast material (e.g., iodine). Example separation filter materials may include rhodium, silver, and tin filters 120.

The spectral imaging system 700 includes a direct conversion pre-Photon Counter Detection (PCD) device 701 positioned on a substrate, such as the glass substrate 225. As in the embodiment of fig. 4, PCD 701 comprises a flat panel imager, such as an amorphous selenium (α -Se) based SWAD, for photon counting at each pixel location. The alpha-Se slab SWAD imager layer 701 converts the received photons into electrical pulses having pulse heights commensurate with the energy levels of the received photons. At each pixel location of the PCD layer, the threshold and count circuit 230 detects the pulse height to distinguish any one of the LE photons and HE photons received at the PCD layer from incident radiation that is not absorbed by (i.e., transmitted through) the object 12. The threshold circuit may include a pulse threshold circuit for comparing the height of the received pulse to a pulse height threshold. At a pixel, both LE and HE counter circuits count the number of photon interactions as an electronic pulse received at that pixel location. Photons received from radiation that exceed a pulse height threshold may be counted as HE photons, and the HE stores a "bin" maintaining a photon count value in a memory store associated with the pixel location. Similarly, at the same pixel location, photons received from radiation below the pulse height threshold may be counted as LE photons, and the LE "bin" maintains an LE photon count value in a memory store associated with that pixel location.

As shown in fig. 7, the data acquisition circuit 98 includes a programmed processor 99 and memory configured to communicate with each other, which can receive count values from each of the LE and HE bins, and process these values to produce respective LE and HE spectral images of the subject.

In an embodiment, the single shot/dual layer detector 700 further includes an indirect conversion flat panel x-ray detector layer 751 formed on the second glass substrate 226 and attached below the substrate 225 using, for example, a radiation transparent adhesive. In an embodiment, the indirect conversion flat panel x-ray detector layer 751 is an Energy Integrated Detector (EID), such as a column (col-) CsI. However, in general, the rear EID detector 751 is a material having an atomic number greater than the atomic number of the material of the first x-ray detector layer 701.

The rear EID included in single shot/dual layer detector 700 includes a scintillating phosphor layer (phosphor screen) for converting x-ray energy photons into light photons that can be sensed by an associated photosensor (photodetector) array circuit 235, which photosensor array circuit 235 is configured to indirectly capture the energy of the light photons from x-rays transmitted through the subject. For example, an EID phosphor layer may include a phosphor crystal that can capture incident x-rays and convert the captured x-rays into light photons. Although not shown, the top surface of the x-ray detector layer 751 can include a reflective layer, wherein the reflective layer can be made of a highly reflective material. For example, the reflective layer may be coated with a layer of white material, such as titanium dioxide. The reflective layer may reflect scattered photons toward the photosensor array 235 so that the photosensor array captures any scattered photons. Thus, in some examples, incident x-rays may not be fully captured by the front detector (e.g., PCD layer) 701 to count all photon interactions. The uncaptured x-rays may pass through the PCD layer 701, and crystals in the phosphor screen of the EID layer 751 may convert the captured x-rays into light photons for detection.

In an embodiment, rear screen 751 may include a scintillating phosphor layer or material, such as a phosphor crystal, that can capture light photons. In some examples, the phosphor layer may be of a powder or particle type (e.g., gdO2S2: tb, caWO4, baFCl: eu). In other examples, the screen phosphor may consist of nano-sized particles, such as quantum dots, rather than micron-sized particles of a typical "standard" screen, such as GdO2S2: tb. In still other examples, the scintillation material may be perovskite-type. The rear detector phosphor screen may emit photons of light in the visible region (e.g., a photon burst).

The post-detector phosphor screen may include a structured scintillation layer. For example, the rear detector phosphor screen may include a scintillating phosphor needle structure that can capture light photons. In some examples, the post-detector phosphor screen may be formed by CsI: tl, vacuum deposition needle structure. In some examples, different types of scintillating materials and combinations of types can be used for the rear screen.

The photosensor array 235 may include a photosensitive storage element and may include a plurality of switching elements (not shown). The second substrate 226 may have a small optical thickness and, in alternative embodiments, may be disposed between the photosensor array 235 and the phosphor layer 751. The photosensitive memory element and the switching element may be disposed on top of the substrate 226. The photosensor array may include a-Si: h n-i-p photodiodes, MIS type or other types. The photosensor array may be sensitive to light incident on the top side and may have low transmittance at wavelengths emitted by the phosphor screen of the EID layer 751. For example, the photosensor array 235 can have a high optical absorbance (above 90%) at the wavelength of light emitted by the screen of layer 751, such that pixel crosstalk and cross-over effects can be reduced. In an example, the substrate 226 may be glass, plastic, or cellulose having a thickness of 700 microns. The photosensor array 235 can capture light photons and can convert the captured light photons into electrical signals, which can be used by data acquisition electronics (separate from the detector 700) to produce a digital image. For example, each switching element may correspond to a pixel of an image such that switching a particular column, row, group of pixels may result in readout of a set of pixel values to produce the image.

In the system of fig. 7, a single shot/dual layer detector 700 combines an alpha-Se PCD front detector and an energy integration CsI back detector, with a thickness of alpha-Se typically 200-300 microns. Transmitted photons from the alpha-Se PCD, predominantly above 33keV, can be integrated by the CsI detector and added to the top energy bin of the alpha-Se PCD to ensure high quantum efficiency and SNR. In an embodiment, the number of photons of the SWAD PCD in high energy is approximately one percent of the total number of photons detected by the SWAD and CsI EID together. Adding CsI EID as a post detector provides substantial improvement in the signal-to-noise ratio (SDNR) of the spectral image.

As shown in fig. 7, the data acquisition circuit 98, including the programmed processor 99, may receive the integrated HE photon energy values from each pixel and post-process the values to enhance the HE spectral image of the subject.

Fig. 8 depicts a further embodiment of a spectral imaging system, including a single shot/dual layer detector 800. This embodiment optionally includes a spectral separation filter 120, or alternatively includes a conventional X-ray spectral imaging filter located near the front of the X-ray tube source 15. As in the embodiment of fig. 7, the spectral imaging system 800 includes a direct conversion pre-Photon Counting Detection (PCD) device 801 located on a substrate (e.g., glass substrate 225). PCD 801 includes a flat panel imager, such as an amorphous selenium (α -Se) based SWAD, for photon counting at each pixel location. The alpha-Se slab SWAD imager layer 801 converts the received photons into electrical pulses having pulse heights commensurate with the energy levels of the received photons. At each pixel location of the PCD layer, a threshold (binning) and counting circuit 230 distinguishes LE photons from received HE photons and stores respective count values of detected LE and HE photons received at the pixel. The threshold and counting circuitry interfaces with the data acquisition circuitry 98, and the data acquisition circuitry 98 includes a programmed processor 99, the programmed processor 99 receiving LE or HE photons stored at each pixel for generating corresponding LE and HE spectral images.

The single shot/dual layer detector 800 embodiment of fig. 8 also includes an indirect conversion flat panel x-ray detector layer 851 formed on the second substrate 226 and attached below the substrate 225 using, for example, an adhesive. In an embodiment, the indirect conversion flat panel x-ray detector layer 851 is an Energy Integrating Detector (EID), such as CsI. In general, the rear EID detector 851 is a material having an atomic number greater than the atomic number of the material of the first x-ray detector layer 801.

As in the embodiment of fig. 7, adding the CsI EID layer 851 as a post detector provides a substantial improvement in the signal-to-noise ratio (SDNR) of the spectral image, as this layer captures transmitted photons from the a-Se PCD, i.e., HE photons, predominantly above 33keV, which can be integrated by the CsI detector and added to the top energy bin of the a-Se PCD to ensure high quantum efficiency and SNR. In particular, the data acquisition circuit 98, including the programmed processor 99, may receive the integrated HE photon energy values from each pixel and post-process the values to enhance the HE spectral image of the subject.

Fig. 9 depicts a further embodiment of a spectral imaging system, including a single shot/dual layer detector 900. The embodiment of fig. 9 comprises a combination of spectral separation filters 120, the spectral separation filters 120 being located in the vicinity of the front of the X-ray radiation source 15 for absorbing a part of the X-ray radiation output of the radiation source 15. The spectral imaging system includes a direct conversion flat panel x-ray front detector 901 located on a substrate (e.g., glass substrate 225). The direct conversion flat panel x-ray front detector 901 is first exposed to the x-ray beam transmitted through the imaging subject. In an embodiment, the direct conversion flat panel x-ray detector layer 901 is an Energy Integrating Detector (EID) including, but not limited to, photoconductive materials such as α -Se. In this embodiment, the α -Se EID flat panel X-ray detector layer 901 absorbs X-ray photons and excites electrons in the photoconductor into a conductive state in the conduction band of the material. Electrons in the conduction band move along the electric field lines in the presence of an electric field. Thus, by applying an electrical potential across the α -Se EID material, the charge released by absorption of X-ray radiation can be collected. The semiconductor pixel circuit 930 located between the front imaging detector 901 and the substrate 225 may include a capacitor array or similar charge storage matrix for storing collected charge released as a result of absorption of X-ray radiation.

In an embodiment, the front detector is of atomic number Z ₁ And thickness T ₁ And a high spatial resolution material selected to preferentially absorb lower energy x-rays while transmitting higher energy, thereby allowing a low energy image to be formed. Its high spatial resolution is used to preserve image detail information, i.e. small structures and sharp edges. Amorphous selenium detector (Z) ₁ ＝34，T ₁ =150 micrometers) is an example of a front detector material.

As further shown in fig. 9, pixel circuits 930 comprising a charge storage matrix located between the front imaging detectors 901 interface with the data acquisition circuit 98, and the data acquisition circuit 98 includes a programmed processor 99 and memory for receiving values associated with LE or HE photons stored at each pixel and processing the data to generate corresponding LE and HE spectral images.

In this embodiment, the exit surface of the front detector is located, and is shown asSandwiched between the front integrating detector 901 and the substrate 225 is a further spectral filter 940 for modulating the x-ray energy spectrum leaving the front detector 901. The filter 940 may include one or more filters having an atomic number Z _2F To Z _NF And the thickness range is T _2F To T _2F A material therebetween. In particular, the material of filter 940 is selected to attenuate low energy photons and facilitate device fabrication, e.g., glass used as a substrate for the active matrix of the pre-fabrication detector. The thickness of the filter 940 is tuned to the desired tradeoff between energy modulation and system sensitivity.

Also included is an indirect conversion flat panel x-ray detector layer 951 formed on the second substrate 226 and attached below the substrate 225 using, for example, an adhesive. In an embodiment, the indirect conversion flat panel x-ray detector layer 951 is an Energy Integrating Detector (EID) of a material such as CsI. The second indirectly converted flat panel x-ray post detector may be of atomic number Z ₃ Greater than atomic number Z of the front detector material ₁ And has a thickness T ₃ . In an embodiment, the atomic number of the post-detector is selected to effectively detect higher energy photons, thereby forming a high energy image. It can be matched to the K-edge of contrast agent, e.g. CsI for iodine: tl to increase the saliency of the contrast agent in the high-energy image. The scintillator screen of the post-detector may be transparent or optically turbid, structured or unstructured (e.g. columnar Cs: tl or powder Gd ₂ O ₂ S: tb), and includes an optically reflective or absorptive backing. The scintillator is coupled to the photodetector or photosensor array 235 by direct deposition, pressure contact, or in some embodiments by using a fiber optic plate to transmit the x-ray induced light image to the photodetector without light diffusion. The light sensor may be a-Si: h-photodiodes, MIS types, or other types known in the art. The Thin Film Transistor (TFT) switching element may be α -Si: type H, metal Oxide (MOTFT) or other types known in the art

As further shown in fig. 9, a data acquisition circuit 98 including a programmed processor 99 may receive the integrated HE photon energy values from each pixel and post-process the values to enhance the HE spectral image of the subject.

Fig. 10 depicts a further embodiment of a spectral imaging system including a dual shot exposure/dual layer detector 1000 with a switchable kVp and filter arrangement system. The embodiment of fig. 10 includes a combination of low energy spectral filters 124, the low energy spectral filters 124 being located near the front of the X-ray radiation source 15 for providing LE radiation fractions, such as 28keV X-ray radiation output of the radiation source 15 during a first exposure, and additionally includes a high energy spectral filter 126, the high energy spectral filter 126 being located near the front of the X-ray radiation source 15 for providing HE radiation fractions, such as 49keV X-ray radiation output of the radiation source 15 during a second exposure.

The dual layer detector 1000 of fig. 10 corresponds to the dual layer detector 900 of fig. 9 and includes a direct conversion flat panel x-ray front detector 1001 located on a substrate such as a glass substrate 225. The direct conversion flat panel x-ray front detector 1001 is first exposed to an x-ray beam transmitted through the imaging subject 12. In an embodiment, the direct conversion flat panel x-ray detector layer 1001 is an Energy Integrating Detector (EID) including, but not limited to, photoconductive materials such as α -Se. In this embodiment, the α -Se EID flat panel X-ray detector layer 1001 absorbs X-ray photons and excites electrons in the photoconductor into a conductive state in the conduction band of the material. Electrons in the conduction band move along the electric field lines in the presence of an electric field. Thus, by applying an electrical potential across the α -Se EID material, the charge released by absorption of X-ray radiation can be collected. The semiconductor pixel circuit 1030, which is located between the front imaging detector 1001 and the substrate 225, may include a capacitor array or charge storage matrix for storing collected charge released as a result of absorbing X-ray radiation.

In an embodiment, the front detector is of atomic number Z ₁ And thickness T ₁ And a high spatial resolution material selected to preferentially absorb lower energy x-rays while transmitting higher energy, thereby allowing a low energy image to be formed. Its high spatial resolution is used to preserve image detail information, i.e. small structures and sharp edges. Amorphous selenium detector (Z) ₁ ＝34，T ₁ =150 microMeter) is an example of a front detector material.

As further shown in fig. 10, pixel circuit 1030, including a charge storage matrix located between the front imaging detectors 1001, interfaces with data acquisition circuit 98, and data acquisition circuit 98 includes a programmed processor 99 for receiving values associated with LE photons stored at each pixel and processing the data to generate an LE spectral image.

In this embodiment, located at the exit surface of the front detector, and shown sandwiched between the front integrating detector 1001 and the substrate 225, is an additional spectral filter 1040 for modulating the x-ray energy spectrum exiting the front detector 901. The filter 1040 may include one or more filters having an atomic number Z _2F To Z _NF And the thickness range is T _2F To T _2F A material therebetween. In particular, the material of filter 1040 is selected to attenuate low energy photons and to facilitate device fabrication, e.g., glass used as a substrate for the active matrix of the pre-fabrication detector. The thickness of filter 1040 is tuned to the desired tradeoff between energy modulation and system sensitivity.

Also included is an indirect conversion flat panel x-ray detector layer 1051 formed on the second substrate 226 and attached below the substrate 225 using, for example, a radiation transparent adhesive. In an embodiment, the indirect conversion flat panel x-ray detector layer 1051 is an Energy Integrating Detector (EID) of a material such as CsI. The second indirectly converted flat panel x-ray post detector may be of atomic number Z ₃ Greater than atomic number Z of the front detector material ₁ And has a thickness T ₃ . In an embodiment, the atomic number of the post-detector is selected to effectively detect higher energy photons, thereby forming a high energy image. It can be matched to the K-edge of contrast agent, e.g. CsI for iodine: t1 to increase the saliency of the contrast agent in the high-energy image. The scintillator screen of the post-detector may be transparent or optically turbid, structured or unstructured (e.g. columnar Cs: tl or powder Gd ₂ O ₂ S: tb), and includes an optically reflective or absorptive backing. Scintillator coupling to photodetection by direct deposition, pressure contact, or in some embodiments by use of fiber optic platesAn array 235 to transmit the x-ray induced light image to a photodetector without light diffusion. The light sensor may be a-Si: h-photodiodes, MIS types, or other types known in the art. The Thin Film Transistor (TFT) switching element may be a-Si: type H, metal Oxide (MOTFT) or other types known in the art

As further shown in fig. 10, a data acquisition circuit 98 including a programmed processor 99 may receive the integrated HE photon energy values from each pixel and post-process the values to generate an HE spectral image of the subject.

In an exemplary operation, in a double shot (double exposure) method, the first exposure is performed with a low energy beam and appropriate filters (e.g., 28keV and Rh), and the second exposure is performed with a higher energy beam and filters (e.g., 49keV and copper). Motion artifacts may be present but may be mitigated by registering the appropriate filter in front of the radiation source in successive time instances using fast keV switching and a rotatable filter wheel. In the simplest case, the image data from the two detector layers are added to form LE and HE images for dual energy subtraction. That is, the results from both the front and rear detectors for the first (28 keV) exposure time add to form an LE image, and the results from the two detectors for the second (49 keV) exposure time add to form an HE image. For a 28keV exposure, the contribution of the CsI layer detector will be small, as more absorption will occur in the front detector Se layer. For an exposure of 49keV, the contribution of CsI will be large, as many higher energy x-rays will penetrate the Se layer. Benefits will include: (i) greater energy separation in LE and HE images, (ii) LE images will look very similar to conventional mammography performed at 28keV, and (iii) because of the CsI layer, the x-ray absorption of HE beam is greater than when Se is used alone. Furthermore, four sets of image data may be useful for multiple material decomposition.

Fig. 11A-11E depict simulation result images 1100 of digital mammography, contrast-enhanced dual-energy mammography, digital breast tomosynthesis, and contrast-enhanced digital breast tomosynthesis. The upper half of each image is from a digital breast phantom (with breast structure) and the lower half is uniform breast tissue with 25% density. A set of iodine objects (each arranged in a 4 x 4 array 1101 and 1102) is inserted into both the upper and lower halves of the phantom. The iodine objects were spheres of diameters 2, 3, 5 and 8mm, and iodine concentrations were 1, 2, 3 and 5mg/ml, respectively.

In the analog image, fig. 11A shows an analog digital mammography image produced using a single layer detector configured with a 50 micron thick Rh spatial separation filter (28 kVp) and a 300 micron thick a-Se detector.

Fig. 11B shows an analog digital mammography image produced by using a single layer detector configured with a 237 micron thick copper filter (49 kVp) and 300 micron thick alpha-Se front detector.

Fig. 11C shows a contrast enhanced dual energy mammography image (e.g., the dual layer detector of fig. 9) configured with a 200 micron thick Rh spatial separation filter (49 kVp), a 200 micron thick alpha-Se front detector, a 700 micron thick glass, and a 400 micron thick post column CsI detector.

Fig. 11D shows a contrast enhanced dual energy mammography image (e.g., the dual layer detector of fig. 9) configured with a 237 micron thick copper spatial separation filter (49 kVp), a 200 micron thick alpha-Se front detector, a 700 micron thick glass substrate, and a 400 micron thick col-CsI back detector.

Fig. 11E shows a contrast enhanced dual energy mammography image (e.g., the dual layer detector of fig. 9) configured with a 200 micron thick Rh spatial separation filter (49 kVp), a 200 micron thick col-CsI front detector, a 700 micron thick glass substrate, and a 400 micron thick col-CsI back detector.

Fig. 12 shows a table 1150 depicting signal-to-noise ratios in the simulated contrast enhanced dual energy mammography results of fig. 11C, 11D, and 11E. As shown in table 1150, an optimal SDNR of 2.83 is achieved for the lower half of the image portion, e.g., image half 1160 of fig. 11C, in contrast to 1.65 for the upper half of image 1170, which is obtained using the analog single shot/dual layer detector of fig. 9 configured with a 200 micron thick Rh spatial separation filter (49 kVp), a 200 micron thick alpha-Se front detector, 700 micron thick glass, and a 400 micron thick post column CsI detector. Also, for the lower half image portions of fig. 11D and 11E, the sdn r values are increased compared to the corresponding sdn r 65 implemented for their corresponding top image portions. As shown, the embodiment of 11C, 11D implementing the dual layer a-Se front detector and CsI back detector performs better than the dual layer embodiment using CsI front detector and back detector.

Fig. 13A-13C depict images specifically illustrating the practical problem of DBT that the dual layer detector scheme of the embodiment of fig. 7 and 9 is intended to overcome. In fig. 13A-13C, which depict LE, HE and CE DBT image slices, respectively, significant patient motion is observed between HE and LE images, and this problem is exacerbated in DBT and CEDBT.

Embodiments of the systems and methods described herein overcome some of the disadvantages of various digital radiography systems and film screen radiography systems by enabling a form of x-ray imaging that extracts material specific information from a volume of interest, such as the location and intensity of contrast agents that have been injected into the body. Furthermore, this is done in a single exposure, eliminating the problem of patient movement between exposures. Several ways of practicing the invention are disclosed, including the use of a dual layer detector and also the use of a photon counting detector. The system and method enable the acquisition of clinically valuable information in a digital breast tomography system, such as 3D breast tissue density, 3D microcalcification distribution and type, 3D distribution of contrast agent (e.g., iodine), and 3D material decomposition of breast lesions (e.g., masses) of interest.

In mammography, spectral tomography provides more complete and accurate diagnostic information about the breast than traditional 2D Full Field Digital Mammography (FFDM), spectral FFDM, and traditional DBT. This information includes 3D breast tissue density, 3D microcalcification distribution and type, 3D distribution of contrast agent (e.g., iodine), and 3D material decomposition of breast lesions of interest (e.g., masses).

Accordingly, the systems and methods disclosed herein provide for performing material-selective breast imaging, which enables the acquisition of clinically valuable information such as the 3D location of contrast agents, while eliminating image artifacts due to patient motion.

The terminology used herein is for the purpose of describing particular embodiments only and is not intended to be limiting of the invention. As used herein, the singular forms "a", "an" and "the" are intended to include the plural forms as well, unless the context clearly indicates otherwise. It will be further understood that the terms "comprises" and/or "comprising," when used in this specification, specify the presence of stated features, integers, steps, operations, elements, and/or components, but do not preclude the presence or addition of one or more other features, integers, steps, operations, elements, components, and/or groups thereof.

In the claims below, the corresponding structures, materials, acts, and equivalents of all means or step plus function elements, if any, are intended to include any structure, material, or act for performing the function in combination with other claimed elements as specifically claimed. The description of the present invention has been presented for purposes of illustration and description, but is not intended to be exhaustive or limited to the invention in the form disclosed. Many modifications and variations will be apparent to those of ordinary skill in the art without departing from the scope and spirit of the invention. The embodiments were chosen and described in order to best explain the principles of the invention and the practical application, and to enable others of ordinary skill in the art to understand the invention for various embodiments with various modifications as are suited to the particular use contemplated.

Claims

1. An apparatus, comprising:

a separation filter for separating a radiation spectrum from the X-ray radiation source into a first energy band and a second energy band for imaging incident radiation on the object;

a substrate;

an x-ray photon counting detector formed on the substrate, the x-ray photon counting detector comprising an array of detector pixels, each detector pixel comprising a sensor for detecting interaction of individual x-ray photons of incident radiation transmitted through the imaging subject during a fixed period of time; and

Each detector pixel of the array having associated counting circuitry operable to generate a first electrical signal representing a respective count of the number of detected interactions of the respective x-ray photons of the first energy level band and a second electrical signal representing a respective count of the number of detected interactions of the respective x-ray photons of the second energy level band, wherein the first and second electrical signals from the detector pixels of the array provide respective energy spectral images of the imaging subject.

2. The apparatus of claim 1, wherein each detector pixel sensor generates an electrical pulse having a height attribute commensurate with an energy level of interacted x-ray photons, each associated counting circuit of the array comprising a pulse height threshold discriminator circuit for increasing a count of detected interactions of individual x-ray photons having a height attribute at or above a particular threshold energy level for discriminating between high and low energy levels.

3. The apparatus of claim 2, wherein the radiation of the first energy level band has a greater energy than the radiation of the second energy level band, the particular threshold energy level corresponding to a pulse height attribute associated with the first energy level band.

4. The apparatus of claim 1, wherein the imaging subject comprises a contrast material having a characteristic K-edge atomic band level, the separation filter absorbing X-ray radiation near the K-edge atomic band level.

5. The apparatus of claim 4, wherein the contrast agent is iodine and the separation filter comprises a material selected from the group consisting of rhodium, silver, palladium, indium, and tin.

6. The apparatus of claim 4, wherein each detector pixel sensor comprises an amorphous selenium (a-Se) based field shaping multi-well avalanche detector (SWAD).

7. The apparatus of claim 6, wherein the x-ray photon counting detector is a front detector formed on a substrate, the apparatus further comprising:

a rear detector located below the substrate, the rear detector comprising:

a scintillation screen for converting incident radiation comprising x-ray photons transmitted through the imaging subject and through a first energy band of the front detector into light photons; and

a photosensor array disposed between the scintillation screen and the substrate, the photosensor array being operable to capture photons of light from the scintillation screen and to convert the captured photons of light into additional electrical signals that are operable to be combined with the first electrical signals from the detector pixel array to obtain an image of the imaging subject.

8. The apparatus of claim 7, wherein the post-detector is an integrating detector and the material of the scintillation screen matches a characteristic K-edge atomic energy band level of the contrast material.

9. The device of claim 7, wherein the flicker screen is of a structured or columnar type, or of an unstructured or granular type.

10. The apparatus of claim 7, wherein the post detector is a columnar CsI energy integrating detector.

11. The apparatus of claim 7, wherein the flicker screen further comprises a backing comprising one of a reflective surface or an absorptive surface.

12. The apparatus of claim 7, wherein the photosensor array comprises:

a plurality of photosensitive storage elements for capturing at least a portion of light photons from the scintillation screen; and

a plurality of switching elements, wherein one of the plurality of switching elements corresponds to one of the plurality of photosensitive storage elements, respectively; a transparent metal bias layer and a transparent 2D patterned metal layer, wherein the transparent 2D patterned metal layer faces the flicker screen.

13. An apparatus, comprising:

A first substrate;

an x-ray photon pre-count detector formed on the first substrate, the x-ray photon count detector comprising an array of detector pixels, each detector pixel comprising a sensor for detecting interaction of individual x-ray photons of incident radiation transmitted through the imaging subject during a fixed period of time;

each detector pixel of the array having associated counting circuitry operable to generate a first electrical signal representing a respective count of the number of detected interactions of individual x-ray photons of a first energy level band and a second electrical signal representing a respective count of the number of detected interactions of individual x-ray photons of a second energy level band, wherein the first and second electrical signals from the detector pixels of the array provide respective low-energy and high-energy spectral images of the imaging subject; and

a rear detector formed on a second substrate and located below the first substrate, the rear detector comprising:

a scintillation screen for converting incident radiation comprising x-ray photons transmitted through an imaging subject and through the second energy band of the front detector into light photons; and

A photosensor array disposed between the scintillation screen and a second substrate for the rear detector, the photosensor array being operable to capture photons of light from the scintillation screen and to convert the captured photons of light into additional electrical signals, the additional electrical signals being operable to be combined with the second electrical signals from the front detector pixel array to obtain an image of the imaging subject.

14. The apparatus of claim 13, wherein each front detector pixel sensor generates an electrical pulse having a height attribute commensurate with an energy level of interacted x-ray photons, each associated counting circuit of the array comprising a pulse height threshold discriminator circuit for increasing a count of detected interactions of individual x-ray photons having a height attribute at or above a particular threshold energy level.

15. The apparatus of claim 14, wherein the radiation of the second energy level band has a greater energy than the radiation of the first energy level band, the particular threshold energy level corresponding to a pulse height attribute associated with the first energy level band.

16. The apparatus of claim 13, wherein the imaging subject comprises a contrast material having a characteristic K-edge atomic band level, the separation filter having an X-ray absorption edge for absorbing X-ray radiation near the K-edge atomic band level.

17. The apparatus of claim 16, wherein the contrast agent is iodine and the separation filter comprises a material selected from the group consisting of rhodium, silver, palladium, indium, and tin.

18. The apparatus of claim 13, wherein each front detector pixel sensor comprises an amorphous selenium (a-Se) based field shaping multi-well avalanche detector (SWAD).

19. The apparatus of claim 16, wherein the post detector is an integrating detector and the material of the scintillation screen matches a characteristic K-edge atomic energy band level of the contrast material.

20. The device of claim 19, wherein the flicker screen is of a structured or columnar type, or of an unstructured or granular type.

21. The apparatus of claim 18, wherein the post detector is a columnar CsI energy integrating detector.

22. The apparatus of claim 13, wherein the flicker screen further comprises a backing comprising one of a reflective surface or an absorptive surface.

23. The apparatus of claim 22, wherein the photosensor array comprises:

a plurality of switching elements, wherein one of the plurality of switching elements corresponds to one of the plurality of photosensitive storage elements, respectively; a metal bias layer and a 2D patterned metal layer, wherein the 2D patterned metal layer faces the flicker screen.

24. An apparatus, comprising:

a separation filter for spectrally separating radiation from the X-ray radiation source into a first energy band and a second energy band for imaging incident radiation on the object, the second energy band having an energy greater than the first energy band;

a glass substrate;

a front-integrating detector formed on the glass substrate, the front-integrating detector comprising a first pixel sensor for directly converting first energy band x-ray photons of incident radiation transmitted through the imaging into a first image signal configurable to form a low-energy image of the imaging subject; and

a post-integration detector formed below the glass substrate, the post-integration detector including a second pixel sensor for indirectly converting second energy band x-ray photons of incident radiation transmitted through the imaging subject and through the pre-integration detector and the substrate into a second image signal configurable to form a high energy image of the imaging subject.

25. The apparatus of claim 24, wherein the front integrating detector comprises:

a first photoconductive layer for converting incident radiation comprising x-ray photons transmitted through a first energy band of the imaging subject into electrical charge; and

a first charge storage array disposed between the first photoconductive layer and the substrate for storing charge associated with the converted x-ray photons.

26. The apparatus of claim 25, wherein the post-integration detector comprises:

a scintillation screen for converting incident radiation comprising x-ray photons transmitted through the imaging subject, the front integrating detector, and the glass substrate at a second energy band into light photons; and

a photosensor array disposed between the scintillation screen and the glass substrate, the photosensor array being operable to capture photons of light from the second scintillation screen and to convert the captured photons of light into a second imaging signal.

27. The apparatus of claim 26, wherein the imaging subject comprises a contrast material having a characteristic K-edge atomic band level, the separation filter having an X-ray absorption edge for absorbing X-ray radiation near the K-edge atomic band level.

28. The apparatus of claim 27, wherein the scintillation screen has a material that matches a characteristic K-edge atomic energy band level of the contrast material.

29. The device of claim 26, wherein the flicker screen is of a structured or columnar type, or of an unstructured or granular type.

30. The apparatus of claim 26, wherein the flicker screen further comprises a backing comprising one of a reflective surface or an absorptive surface.

31. The apparatus of claim 26, wherein the photosensor array comprises:

a plurality of switching elements, wherein one of the plurality of switching elements corresponds to one of the plurality of photosensitive storage elements, respectively; a metal bias layer and a transparent 2D patterned metal layer, wherein the transparent 2D patterned metal layer faces the flicker screen.

32. The apparatus of claim 24, further comprising:

a filter disposed at the exit surface of the front detector, the filter comprising a material operable to modulate X-ray radiation exiting the front detector and attenuate X-ray photons of the first energy band.

33. The apparatus of claim 24, wherein the first photoconductive layer of the front energy integrating detector comprises amorphous selenium material and the scintillation screen of the rear energy integrating detector comprises CsI material.

34. An x-ray imaging system, comprising:

an x-ray radiation source;

a first low energy band filter configured for registration in front of the x-ray radiation source to permit transmission of first low energy band x-ray radiation at a first instance in time for incidence on an imaging subject;

A second high energy band filter configured for subsequent registration in front of the x-ray radiation source to permit transmission of second high energy band x-ray radiation at a second time instance for incidence on the imaging subject; and

a dual layer x-ray radiation detector comprising:

a glass substrate;

a front energy integrating detector formed on the glass substrate, the front energy integrating detector comprising a first pixel sensor for directly converting photons of a first low energy band of incident x-ray radiation transmitted through the imaging subject during the first time instance into a first image signal and photons of a second higher energy band of incident x-ray radiation transmitted through the imaging subject during the second time instance into a second image signal; and

a rear energy integrating detector formed below the glass substrate, the rear energy integrating detector comprising a second pixel sensor for indirectly converting photons of a first low energy band of incident x-ray radiation transmitted through the imaging subject and through the front energy integrating detector and the glass substrate during the first time instance into additional first image signals, and photons of a second higher energy band of incident x-ray radiation transmitted through the imaging subject and through the front energy integrating detector and the glass substrate during the second time instance into additional second image signals;

Wherein the first image signal from the front detector and the further first image signal from the rear detector are combinable to form a low energy image of the imaging subject, and

the second image signal from the front detector and the further second image signal from the rear detector may be combined to form a high energy image of the imaging subject.

35. The system of claim 34, wherein the post energy integration detector comprises:

a scintillation screen for converting incident radiation comprising x-ray photons transmitted through the imaging subject and through the first low energy band and the second higher energy band of the front energy integrating detector into light photons; and

a photosensor array disposed between the scintillation screen and the glass substrate, the photosensor array being operable to capture photons of light from the scintillation screen and to convert the captured photons of light into respective ones of the additional first image signals and the additional second image signals.

36. The system of claim 34, wherein the front energy integrating detector comprises:

a first photoconductive layer for converting incident radiation comprising x-ray photons transmitted through a first low energy band and a second higher energy band of the imaging subject into corresponding charges; and

A charge storage array disposed between the first photoconductive layer and the substrate for storing respective charges associated with converted x-ray photons for forming respective ones of the first image signal and the second image signal.

37. A method for x-ray imaging, comprising:

providing a dual layer x-ray radiation detector below an object to be imaged with incident x-ray radiation, the dual layer x-ray radiation detector comprising:

a glass substrate;

a front energy integration detector formed on the glass substrate, including a first pixel sensor; and

a rear energy integrating detector formed below the glass substrate, including a second pixel sensor;

registering a first low energy band filter in front of the x-ray radiation source to allow transmission of first low energy band x-ray radiation at a first time instance for incidence on the imaging subject;

registering a second high energy band filter in front of the x-ray radiation source to permit transmission of second high energy band x-ray radiation for incidence on the imaging subject at a second time instance, and a dual layer x-ray radiation detector to receive the first low energy band x-ray radiation that passed through the imaging subject at the first time instance and second high energy band x-ray radiation that passed through the imaging subject at the second time instance;

Using the first pixel sensor of the front energy integrating detector, directly converting photons of a first low energy band of incident x-ray radiation transmitted through the imaging subject during the first time instance into a first image signal and photons of a second higher energy band of incident x-ray radiation transmitted through the imaging subject during the second time instance into a second image signal;

using the second pixel sensor of the rear energy integrating detector, indirectly converting photons of a first lower energy band of incident x-ray radiation transmitted through the imaging subject and through the front energy integrating detector and the glass substrate during the first time instance into additional first image signals, and indirectly converting photons of a second higher energy band of incident x-ray radiation transmitted through the imaging subject and through the front energy integrating detector and the glass substrate during the second time instance into additional second image signals;

forming a low energy image of the imaging subject from the first image signal and the further first image signal,

a high energy image of the imaging subject is formed from the second image signal and the further second image signal.